As is known, positron-emission tomography (PET) is widely used, i.e., in the medical sector for the detection of structures and formations in human tissues, and consists in supplying a tracer isotope bound to an active molecule to tissues to be examined. The decay of the isotope in the tissues causes generation of positrons that are annihilated with electrons and give rise to gamma rays, which can be detected by a layer of inorganic crystals (scintillator) where a flash is created, which is in turn detected by a photomultiplication structure arranged adjacent to the layer of inorganic crystals.
For example, FIG. 1 shows a typical detector ring 1 present in a PET machine. As may be noted, the detector ring 1 is formed by a plurality of detector blocks 2, circumferentially arranged around the ring, each with radial extension. Referring to FIG. 2, each detector block 2 includes an array of inorganic crystals or scintillators 3 adjacent to an array of photomultipliers 4.
Traditionally, the photomultipliers 4 are formed by photomultiplier tubes (PMTS); however, in the last few years, the use of silicon photomultipliers (SiPms) has been proposed, thanks to the high efficiency that can be obtained (see, for example, “Silicon Photo-multipliers as Photon Detector for PET”, by R. Pestotnik et al., 2008 IEEE Nuclear Science Symposium Conference Record, which is incorporated by reference).
The proposed SiPms are formed as arrays of individual photomultiplier cells, each made up of a plurality of elements for detecting individual photons, typically Geiger-mode avalanche diodes (GMAPs), made, for example, as described in US 2009-0184317 and US 2009-0184384, which are incorporated by reference. In particular, the avalanche diodes operate at reverse biasing voltages that are a few volts higher than the breakdown voltage, and each avalanche diode detects an individual photon. In fact, as shown in FIG. 3, each avalanche diode 5 is coupled to the supply Vb through a respective quenching resistor 6, disposed (e.g., integrated) in series, and forms with the latter a pixel 7. In each pixel 7, the quenching resistor 6 is able to quench the avalanche current and reset only the relevant avalanche diode 5 after detection of a photon. In a photomultiplier cell 8, the pixels 7 are coupled in parallel to one another so that the currents detected by each of the individual pixels 7 are added together. The intensity of the total current of the photomultiplier cell 8 is thus given by the analog superposition of the signals (binary signals in an embodiment) produced by all activated pixels 7, which is, in turn, proportional to the number of incident photons (to a first approximation, if multiple hits on different pixels 7 are neglected).
Thus, in case of weak flows of photons, a photomultiplier cell 8 of SiPm diodes behaves as an analog or proportional device, whereas the individual pixels operate in digital or Geiger mode. For this reason, silicon photomultipliers are also frequently represented as digital-to-analog conversion devices.
Currently, in general, each photomultiplier cell is manufactured in a chip with dimensions of approximately 4×4 mm2, or, at the most, approximately 5×5 mm2, and the various chips are arranged near each other so as to form more extensive detection surfaces. In particular, the chips are bonded to an intermediate substrate (generally a die of semiconductor material), which is, in turn, bonded to a base substrate, which is larger and is generally obtained from a printed circuit board (PCB) or from PCB material. The group formed by the base substrate, the intermediate substrate, and the array of cells is then bonded to the array of scintillators 3, for example, with the interposition of an optical grease (see, for example, “Evaluation of Arrays of Silicon Photomultipliers for Beta Imaging”, E. Heckathorne, L. Tiefer, F. Daghighian, M. Dahlborm, 2008 IEEE Nuclear Science Symposium Conference Record, which is incorporated by reference).
However, arranging the cells near each other typically requires the presence of free areas for providing the connections (paths, pads for the connection wires, etc.). In at least some known devices, for example, the active areas represent approximately 50-60% of the total area. Consequently, there are “dead” or “blind” areas with no pixels, where, consequently, photons are lost. This entails a reduction in the theoretical resolution of the PET, and may entail a lengthening of the times for carrying out the medical examination.
This problem may increase when the single chips that integrate a cell are bonded to an own intermediate substrate, on account of the tolerance of bonding the intermediate substrates.
Attempts at directly bonding the chips to the base substrate have not solved the problem since the structure tends to undergo deformation and to lose planarity (i.e., warp), because of the large dimensions of the base substrate, of the layer of cement used, and of the bonding process, thus preventing proper contact between the photomultipliers and the photomultiplier cells, and thus worsening the detection efficiency.
It is also possible to reduce the dead spaces by providing through silicon vias (TSVs) through the base substrate and through the intermediate substrate, when present, but this solution typically does not solve the problem of warpage, and may be costly.
On the other hand, the creation of the array of cells in a single integrated device, for example, of approximately 32×32 mm2, gives rise to problems of yield, and also devices of these dimensions tend to undergo considerable deformation due to stress. In addition, also in this case it would probably be necessary to create through vias through the base substrate, with all the problems highlighted above.
On the other hand, the lack of an effective protection of the chips is a problem since bonding the array of photomultiplier cells to the array of scintillators is generally performed in a plant remote from the site where the photomultiplier devices are produced.